1. Field of the Invention
The present invention relates to liquid flows and, in particular, liquid flows at relatively slow speeds within relatively small bores such as those of microchannel structures or microcapillaries. The invention further relates to a liquid outlet link assembly for such liquid flows and to a microchannel structure assembly for microfluidic systems.
2. Background of the Invention
In recent years, microminiaturized fluidic systems are used extensively in analytical chemistry, drug discovery and life sciences. Microchannel structures of 20-100 xcexcm in cross section are used in chemistry to achieve fast speed electrophoretic separation for chromatography. These structures, often referred to as bioelectronic chips or bio chips, could offer a convenient method of isolating, lysing, and detecting of micro-organisms in complex samples and could have applications in drug discovery, genetic testing and separation sciences (e.g. capillary electrophoresis).
For chemistry and drug discovery microfluidic reactors, microchannel structures and microcomponents present the possibility of decreasing the time of technological processes by integrating several units in relatively small areas. This could facilitate the execution of a number of reactions in parallel and increase efficiency in high throughput screening and combinatorial analysis. Overall, sample volumes within microcomponents are in the microliters range, which can save considerably on the amounts of reagents per reaction. For successful control of microreactors, efficient detection and separation of chemical specimens, it is absolutely necessary to provide fluid delivery at the low flow rates.
In the area of life sciences, the study and manipulation of single biological cells on microfluidic structures could potentially enable considerable experimental progress Manipulation of cells is becoming important for clinical diagnostics and genetic measurements. Lysing of cells, cell-cell interaction, interaction and manipulation of a single cells is now possible and it is more efficient using microchannel structures. It is expected that integrated analytical systems will allow genetic measurements and drug screening at the single cell level.
In these wide areas of application of microstructures, pumping systems play a significant role. Delivering required solutions to the sites of reaction, mixing different fluids, creating gradients of concentration of the reagents, controlling the positions of biological samples, transporting and manipulating them are all tasks, which require a highly accurate pumping system. Despite a major effort in developing pumping systems for a microchannel structure, the problem still remains. Many conventionally used pumping systems are operating with significantly bigger volumes of fluids, therefore they cannot provide pumping accuracy or in some cases adequate pumping speed when it comes to establishing flows inside the microstructures with a microchannel diameter from 5 to 100 xcexcm.
Various constructions of positive displacement pumps, including syringe pumps, positive pressure infusion pumps and peristaltic pumps have been used with capillaries. These are, for example, described in U.S. Pat. No. 4,715,786 (Wolff et al. Syringe pumps with microflow rate capabilities to provide precise and reproducible volumetric flow ranges of the order of 0.1 xcexcl to 1 ml per minute have been described, for example, in U.S. Pat. No. 5,630,706 (Yang) and U.S. Pat. No. 5,656,034 (Kochersperger et al). One of the main objects of these inventions has been to deliver pulse free flow, the problem being that the pressure of the fluid inside the syringe pump changes during the stroke of the syringe pump, which stroke is usually controlled by a stepper motor. Unfortunately, such an operation results in a large pressure surge which alters the volumetric flow rate. For example, Japanese Patent Specification No. 4058074A (Nagataka et al) describes a method to reduce fluctuations of the flow in a syringe pump to provide a more stable flow rate by setting the syringe vertically and forming a gas layer between the front surface of the piston forming the syringe pump and the liquid being pumped. This invention, however, is directed towards relatively large flow rates of the order of microliters per minute and would be useful for drug infusion but would not be particularly suitable for microchannel structures and the like, where the flow rates are, as mentioned already, substantially less.
U.S. Pat. No. 4,137,913 (Georgi) describes a method of controlling the flow rate by changing the stroke periods. U.S. Pat. No. 5,242,408 (Jhuboo et al) describes a method of controlling pressure inside a syringe pump by measuring the force acting on the plunger and detecting an occlusion. Unfortunately, heretofore, such syringe and positive displacement pumps are relatively inefficient at delivering fluid flow at rates of the order of nanoliters per minute, which flow rate is required to transport liquids in microchannel structures. Generally, the limitation on the flow rate is the movement accuracy of the various mechanical parts of the syringe pump such as the stepper motor, plunger, valves, and so on. However, syringe pumps used in high pressure liquid chromatography (HPLC) have achieved volumetric flow rates as low as 0.1 xcexcl/min. A typical example of this is described in U.S. Pat. No. 5,630,706 (Yang). However, for commercially available syringe pumps, the linear displacement of the piston or plunger would be several micrometers per step of the motor controlling the pump. Thus, general sealing surface wear makes it impossible to achieve accuracy for shorter displacements.
A further disadvantage of the syringe pump when used for pumping fluids in microchannel structures, is that it cannot deliver a sufficiently low pumping speed for many applications of the structures.
Typically, a syringe pump would dispense 0.6 xcexcl/min for one step of the motor which then has to be delivered into a microchannel structure possibly having a cross sectional diameter of the order of 40 xcexcm which translates into 1.9 mm/sec. through the microchannel structure which is much too fast for the observation of biological specimens, detection of proteins, single cells and the creation of low gradients of reagents, which is required in many microfluidic applications. Indeed, one can readily appreciate that at this speed, visual observation is difficult and further would not allow for the manipulation or sensing of biological samples. Thus, heretofore, positive displacements pumps and in particular, syringe pumps, while very attractive for their simplicity, have not as of yet been useful for these applications.
Electrokinetic pumps have been proposed for such pumping operations. Pumps based on electroosmotic phenomena have been described in U.S. Pat. No. 3,923,426 (Theeuwes et al) and U.S. Pat. No. 5,779,868 (Wallace Parce et al). When a buffer is placed inside a capillary, the inner surface of the capillary acquires a charge. This is due to the ionisation of the wall or adsorption of ions from the buffer. In the case of silicate glass, the surface silanol groups (Sixe2x80x94OH) are ionised to silanoate groups (Sixe2x80x94Oxe2x88x92). These negatively charged groups attract positively charged cations from the buffer, which form an inner layer of cations at the capillary wall. These cations are not in sufficient density to neutralize all the negative charges, therefore a second layer of cations forms. The inner layer of cations, strongly held by the silanoate groups, forms a fixed layer. The second layer of cations is less strongly held because it is further away from the negative charges, threfore it forms a mobile layer. When an electric field is applied, the mobile layer is pulled toward the cathode. Since ions are in solution, they drag the whole buffer solution with them and cause electroosmotic flow. The distribution of charges due to the formation of charged layers create a potential termed the zeta potential.
This method, originally used for capillary electrophoresis, is recently being used for fluid transport in microstructures and for high speed chromatography in microfluidic chips. However, it still has a number of disadvantages.
The distribution of charges and formation of layers depends on the initial charge of the inner surface of the capillary, which is different for various materials and solutions used. Moreover, it can be reliant on the pH history of the capillary. This makes the control of the zeta potential and therefore electroosmotic flow control a complicated task. The prior art evidences a number of ways to treat the capillary in order to achieve a reproducible flow rate. They indicate that coating the microcapillary with a monomolecular layer of non-cross-linked polyacrylamide can derivatize inner surfaces of a capillary. This coating enhances the osmotic effect and suppresses adsorption of solutes on the walls of the capillary. Others have taught that altering the buffer pH, the concentration of the buffer, the addition of surface-active components, such as surfactants, glycerol, etc. or adding various organic modifiers to the buffer solution may alter electroosmotic flow. In some cases this alteration can cause a reverse of electroosmotic flow or its complete cancellation.
Transport of particles in electroosmotic pumping systems is also difficult, due to the fact that during transport they can acquire an electrical charge and can be moved by the electric field, which in some cases causes flow to reverse.
According to the theory, the mobile layer drags the fluid. As a result electroosmotic flow has a relatively flat flow profile i.e. the flow velocity is rather uniform across the capillary. When a static pressure is opposed to the electroosmotic flow, the resulting flow can produce a turbulence, which doesn""t allow controllable mixing of fluids and biological samples and decreases speed of electroosmotic flow.
For example, U.S. Pat. No. 4,908,112 (Pace et al) suggests the use of electro-osmotic pumps to move fluids through channels less than 100 microns in diameter. A plurality of electrodes was incorporated in the channels, which were etched into a silicon wafer. An electric field of about 250 volts/cm was required to move the fluid to be tested along the channel. However, when the channel is long, a large voltage needs to be applied to it, which may be impractical for highly integrated structures. This US patent specification suggests that the electrodes be staggered to overcome this problem, so that only small voltages could be applied to a plurality of electrodes. However, this requires careful placement and alignment of a plurality of electrodes along the channel.
Electrohydrodynamic (EHD) pumping of fluids is also known and may be applied to small capillary channels. The principle of pumping here is different from electroosmosis. When a voltage is applied, electrodes in contact with the fluid transfer charge to or from the fluid, such that fluid flow occurs in the direction from the charging electrode to the oppositely charged electrode. Electrohydrodynamic (EHD) pumps can be used for pumping resistive fluids such as organic solvents. U.S. Pat. No. 5,632,876 (Zanzucchi Peter John et al) describes the use of both electroosmotic and electrohydrodynamic fluid movement method to establish flow in microcapillaries for polar and non-polar fluids.
One of common problems that is usually encountered in these two types of fluid pumping system is the appearance of gas bubbles, which are easily obtained during pumping as a result of electrolysis. They normally interfere with particle transport, blocking microstructures; requiring additional pressure difference to transport them. Pumping of fluids by pumps based on electroosmosis and electrohydrodynamic phenomena relies on the electrical contact throughout the fluid, which disappears in the presence of bubbles rendering pumping by these methods difficult.
Another method of fluid transport in a microfluidic structure is by mechanical micropumps and valves incorporated within the structure such as described in U.S. Pat. No. 5,224,843 (Van Lintel), U.S. Pat. No. 5,759,014 (Van Lintel) and U.S. Pat. No. 5,171,132 (Miyazaki et al).
As described in U.S. Pat. No. 5,759,014 (Van Lintel), the operation of these pumps is greatly influenced by the compressibility of the fluid and the presence of an air bubble inside the pumping chamber. The pumping speed decreases in the presence of a significant air bubble, sometimes even reducing to zero. Procedures of priming these pumps is complicated and requires a vacuum pump or special injection devices, to prevent appearance of bubbles in the micropumps main pumping chamber. Therefore, it is also impractical to use micropumps as a part of disposable microfluidic biochips.
Another method of pumping fluids in microchannel systems is based on centrifugal force caused by rotation of the microchannel structures at desired speed. In a most common embodiment, the microchannel structure is a disk in a format similar to that of a CD platform. The fluid in this case flows from the centre of rotation to the periphery. Due to opposing surface tension and centrifugal forces at the interface between the fluid medium and air, it is possible to implement an rpm-dependent valves and switches. Therefore this method provides a way to facilitate sequential reactions on chip platform. In U.S. Pat. No. 6,063,589 (Kellogg Gregory et al), the microsystem platforms are described as having microfluidics components, resistive heating elements, temperature sensing elements, mixing structures, capillary and sacrificial valves, as well as methods for using these microsystem platforms for performing biological, enzymatic, immunological and chemical assays. A rotor with a slip ring capable of transferring electrical signals to and from the microsystem platforms is also described in the invention.
While such centrifugal pumps can provide required flow rates in microfluidic systems and integrate components on a single platform, this method has a number of shortcomings. The fluids can only be transported in one direction and no reversed flow is possible. Control of the flow rate in the individual channels is not possible dynamically, but only by designing a specific geometry of the microchannel structures. Therefore mixing is only possible with predefined ratios. Replacement of one of the fluids for a fluid with a different viscosity requires a change in the design of the structure. The valves can only operate only ones when interface between fluid and air presents inside the valve. For the complicated interconnected channel geometries during the filling process air bubbles may appear in some places. This would require an additional increase in the rotation to pump them and therefore would lead to non-reliable experiments particularly in the case of sequentially executed experiments. This is opposed to pressure pumps where multiple pumps can facilitate a filling process individually for each channel, if required. When a microfluidic structure is rotated at a high speed it becomes impossible to visually observe biological samples, which is very important for a number of applications, for example for the study of cellular responses.
Despite several types of pump methods proposed for pumping fluids in the microchannel structures, there is no simple solution, which can be used in many of the applications utilizing microchannel structures. All methods have some disadvantages, which are more or less significant for different applications. For example, it""s not practical to use micromachined pumps in applications of disposable biochips. Integration micromachined pumps with a disposable device would increase the cost of it. In the same example electroosmotic pumps cannot provide a great degree of reliability. It seems to be impractical when every disposable chip needs to be treated before experiment in order to successfully control electroosmotic velocity.
The present invention is directed towards providing a pumping system and method for pumping liquids in microchannel structures to enable an accurate control of flow for flow rates ranging from 100 picoliters per minute to 10 microliters per minute. Such a system should be suitable for pumping both conductive and non-conductive liquids and in particular, for pumping liquids with different viscosities and liquids which contain particles with sizes comparable to the microchannel""s diameter. Thus, such a pumping system should be suitable for delivering liquids with biological samples.
The invention is also directed towards providing a pumping system and method for pumping liquids which will enable accurate mixing of flows in different microchannel structures in a wide range of concentrations, and in particular to the accurate control of liquid gradients in the microchannels. Further, the invention is directed towards providing a pumping system and a method for pumping liquids in microchannel structures which can be accurately controlled either by an operator or in response to some condition of the liquid such as, for example, a speed of reaction or indeed some other phenomenon.
The invention provides a liquid outlet link assembly to provide a steady liquid delivery output rate below 10 xcexcl per minute. This liquid output rate is for delivery through a liquid outlet from a positive displacement pump. Generally, such a positive displacement pump has an immediate step pumping rate which is substantially larger than the delivery rate through the liquid outlet means. Thus, the liquid outlet means has greatly reduced delivery compared to that of the positive displacement pump. The invention provides this liquid outlet link assembly which is a body with a resistance of flow therethrough substantially less than through the liquid outlet means which can be anything from a microcapillary or any such similar system. There is a liquid inlet in the body of the link assembly for connection to the pump and an outlet in the body for connection to the liquid outlet means and a pressure activated expansion means in the body to create a liquid pressure at the liquid outlet to provide the desired liquid delivery flow through the liquid outlet means. Generally, the expansion means comprises a gas bubble and more likely a bubble of air. The volume of the bubble is many multiples of the liquid dispensed in one step of the pump. There can be more than one bubble and indeed any air within the liquid link assembly assists in its operation. It is envisaged that the pressure activated expansion means, instead of being a gas or air bubble, could be an elastic membrane forming part of the body member of the link assembly. It could even be that there is expandable tubing in the link assembly. As mentioned above, ideally the liquid outlet means comprises an elongate microchannel structure in which case it is important to have a liquid pressure which is sufficient to provide the necessary liquid pressure gradient between the proximal end of the microchannel structure and the distal end forming the output of the microchannel structure.
Obviously, the invention provides various forms of control, either means for sensing the flow conditions to ensure that the pump operates in response to the sensed flow conditions and this could be any form of optical flow monitoring. It could equally be a pressure sensing means. Further, the invention provides a pump assembly incorporating the liquid link assembly as described above. The pump can be a syringe pump and generally speaking, the volume pump for one step of the syringe pump is greater than 0.1 xcexcl but could be of the order of 0.2 xcexcl. It is envisaged that more than one syringe pump may be provided and indeed ideally they dispense different volumes, one of the pumps dispensing a volume many multiples of that of the other pump. An electrokinetic pump can be provided and ideally such an electric pump is an electroosmotic pump or could be an electrohydrodynamic pump.
Further, the invention provides a microstructure assembly which has an internal bore of less than 1000 xcexcm2 cross-sectional area. Ideally, there is provided a positive displacement pump operating in which each step operation of the pump dispenses a volume of the order of 0.01 xcexcl. Then again, a liquid link assembly as described above may be used.
It is envisaged that in any liquid link assembly, adjacent the liquid outlet link assembly, there is provided a flow balancing conduit, the cross-sectional area of the body adjacent the inlet substantially equaling the aggregate cross-sectional area of the microchannel structure and the balancing conduit. It can include a liquid take-off means whereby the flow rates of liquid in a liquid link assembly and the liquid outlet means are substantially equal. Thus, where in a microchannel structure, you might have a flow rate normally many multiples of that required in the liquid link assembly to allow the requisite liquid to be delivered through the microchannel structure, by using this flow balancing conduit or other means, the flow rates of both will be substantially the same. For example, the liquid outlet of the body can include a recirculation pipe connected between the liquid outlet and the body of the liquid link assembly.